The Positron Emission Tomography (PET) is a diagnostic technique that obtains images that show the metabolism and functioning of tissues and organs (for example, the central nervous system).
Like other diagnostic techniques in Nuclear Medicine, PET is based in detecting and analyzing the distribution inside the body of radioisotopes which have previously been administered to a patient. The radioisotopes may be taken in orally, they may be inhaled as gas or may be administered through an injection.
Several positron-emitting radioisotopes for medical use are known. The most commonly used is Fluorine-18, which is capable of joining a glucose tracer to get 18-fluoro-deoxy-glucose (18F-FDG). In this way, glucose that is detectable by the emission of radioactive signal is obtained.
After administration of the radioisotopes, the radioisotopes spread throughout the area of the body to be examined and tend to be taken up by e.g. cancer cells. When the radioisotope decays, it emits a positron which after a few millimetres annihilates with an electron. This produces a pair of gamma ray photons moving in opposite direction, each photon having an energy of 511 keV. This pair of gamma ray photons can be detected using a so-called PET scanner. Using the location of detection of both gamma ray photons, the Line of Response (LOR) (which is the line connecting the two locations of detection of the gamma photons) can be reconstructed. This procedure is schematically illustrated in FIG. 14.
FIG. 14 shows a conventional PET scanner 1, in which a bed 3 is provided. Upon this bed, a body 2 of a human or an animal is schematically indicated. Around the circumference of the PET scanner, a plurality of detectors 4 is provided. The gamma ray photons which move in opposite direction are detected respectively by detector 4a and detector 4b. Using this detection, the LOR can be reconstructed.
After collecting several such events, points where multiple LOR's intersect can be determined. These points indicate a concentration of the radioisotope and therefore the possible presence of cancer cells. The PET scanner is coupled to a computer, which is responsible for measuring the amount of radioisotopes absorbed by the body, and determining the LOR's. This way, it is possible to obtain images that provide details of both the structure and function of internal organs and other parts of the body.
In a typical PET protocol, the patient is injected with between 300 to 500 MBq of 18F-FDG (Fluorodeoxyglucose). After allowing one to one and a half hours for uptake, the patient is placed in the scanner for the scan. A typical PET scan on a conventional PET scanner requires around 30 minutes of scanner time.
PET plays an important role in tumour diagnostic. Its accuracy overtakes that of the conventional diagnostic imaging systems as one can see in the following table (from the Journal of Nuclear Medicine Supplement, Volume 42, Number 5, May 2001 and UCLA):
Diagnostic AccuracyConventional Cancer TypeImaging PETBreast67%89%Colorectal80%94%Gastro-Esophageal68%83%Head & Neck65%87%Liver81%93%Lung68%82%Lymphoma64%88%Melanoma80%91%Pancreatic65%81%Testicular68%92%Uterine/Cervical 43%87%
A PET scanner comprises a plurality of detectors. Nowadays, the current best detector for PET is based on LSO (Lutetium Oxyorthosilicate) crystals with a typical size of 4 mm×4 mm×10 mm. The crystals emit light flashes when hit by the gamma photons. These light flashes can be detected using a photomultiplier tube (PMT) which is coupled to the crystal. This has also been schematically indicated in FIG. 14. Detector 4c comprises a segmented crystal 5 and a plurality of PMT's 6. To the person skilled in the art, it will be clear that the plurality of PMT's 6 can also be replaced by a single position sensitive photomultiplier (PSPMT).
The light yield of LSO crystals for 511 keV gamma is about 4000 phe. The full width at half maximum (FWHM) that can be achieved with LSO crystals at 511 keV is around 10%. This limited energy resolution will reduce the ability to remove scattered events, which are one type of noise in the reconstructed image. The typical length (in radial direction, when used in a PET scanner) of the LSO crystal is about 10 mm and this implies that the detector's intrinsic uncertainty of the impact point in the radial direction is around 3 mm, which will lead to an error in the projection of the Line of Response. This can be readily understood when looking at FIG. 14. The PMT which registers an event essentially gives a two-dimensional coordinate. The radial position of where the gamma photon hits the crystal is lost. This loss of information gives rise to a parallax effect, which can lead to an error in the projection of the LOR. This error naturally deteriorates the quality of the reconstructed image.
Another drawback is the shape of the crystals of the detector, which is rectangular parallelepiped and when forming a cylindrical shape from such components (for example, for a PET scanner), it is inevitable to have cracks at the contact points of the crystals.
Another example of medical diagnostic imaging apparatus is a gamma camera. Gamma cameras are also widely used in nuclear medicine. It consists of a single detector plane, formed by a plurality of scintillating crystals, and a collimator in front of it. Only photons in a small angular range reach the detector through holes in the collimator; the others are absorbed by the collimator. Therefore, a two dimensional projection of the source distribution is recorded by the detector plane.
Single Photon Emission Computed Tomography (SPECT) is a nuclear medicine tomographic imaging technique using a gamma camera. SPECT imaging is performed by using the gamma camera for acquiring multiple two dimensional images at different angles. A computer is then used to apply a tomographic reconstruction algorithm to the multiple projections, yielding a three-dimensional image.
Compton cameras are another example of medical diagnostic imaging apparatus. Compton cameras are used to reconstruct gamma-ray emitting radioisotope distributions. The range of applications is vast: apart from diagnosis in nuclear medicine, they may also be used for monitoring of decommissioning of nuclear power plants and also find application in homeland security (e.g. for scanning shipping containers for the presence of radioactive material).
A Compton camera has two detection planes. Photons emitted from the source are scattered in the first plane (Compton scattering) and absorbed in the second plane (photoelectric effect). In both planes the position of the interaction and the energy deposited are measured. The first plane is usually made from semiconductor material and the second plane from scintillating crystals. The detectors are operated in coincidence, so that only photons that interact with both detectors and deposit a total energy within a given window are recorded. Using the location of detection and the energy of the photon, the point of origin of the photon can be calculated, using the so-called Compton formula.
In summary, in nuclear medical imaging, several techniques and apparatus are known (PET, SPECT, gamma camera, Compton camera) which conventionally use detectors based on scintillating crystals.
The detectors based on scintillating crystals suffer from various disadvantages: In PET scanners, cracks are inherently present at the contact points of adjacent parallelepiped crystals. As has been described before for PET scanners, a parallax effect may occur. The crystals used in PET scanners, gamma cameras or Compton cameras have a size of 4 mm×4 mm×10 mm, so this determines the intrinsic error (and the intrinsic spatial resolution) that they have. It is not possible to see anything smaller than 4-5 mm in case of PET scanning. For SPECT, this is in the order of 15-20 mm. The spatial resolution will even be considerably worse since the detectors do not work with 100% efficiency.
Another problem with scintillating crystals is that the Detection Quantum Efficiency (DQE) of scintillating crystals is rather poor. In order to improve this DQE, one would have to increase the length of the crystal to increase the probability of capturing the gamma photons. However, adding more material would not necessarily improve the quality of the signal, since a part of the light produced in the crystals at a location that is relatively far away from the photomultiplier would be attenuated before reaching the PMT. A gamma photon captured by the crystal close to the PMT will yield more photons than a gamma photon captured further away from the PMT. Hence, the energy resolution gets worse at the expense of detecting more gamma photons. Therefore the standard length of scintillating crystals (10 mm) used in gamma detectors is a compromise to obtain reasonable DQE and simultaneously have an acceptable energy resolution.
To overcome the inconveniences of scintillating crystals mentioned before, it has been suggested to use pixelated room temperature solid-state detectors in PET detectors. With pixelated solid-state detectors, high spatial resolution can be achieved due to the fact that the detector can be segmented to the sub-millimetre pixels (or voxels).
One of the problems of using solid-state detectors in a PET scanner is that thick detectors are needed to achieve high gamma ray absorption. In particular, one needs a thickness of 4 cm (when using CdTe) to capture 90% of photons with energy of 511 keV.
In literature, it has been suggested to use very large sensors (e.g. 10 mm×10 mm×10 mm) of Cd(Zn)Te in which a pixel readout chip is coupled to the backside of the crystal. This solution looks easy on paper but in reality it is not. Firstly, the cost of such large CdTe detectors of good quality is very high. Secondly, the time collection of the signal will be very long and hence it will not be possible to use it in PET as trigger. Thirdly, with such a thick detector (10 mm), the energy spectroscopy deteriorates significantly due to trapping and the lifetime of the electron-hole.
U.S. 2007/0057191 discloses a radiological imaging system comprising a first and second imaging apparatus. Said first imaging apparatus comprises an array of (non-pixelated) semiconductor radiation detectors. The array of detectors is connected through wiring to ASICs. Since the detectors used in this system are not pixelated, the accuracy that may be achieved is inherently limited. The detectors used are at least 2 mm×2 mm. The accuracy is thus limited to these dimensions. Furthermore, in practice it will be very complicated and costly to use detectors of these dimensions, so that bigger detectors will be needed. Also, between all of the detectors, gaps are present. These gaps, which represent dead areas in which no events can be detected, and the way the arrays of detectors are packaged and arranged lead to limited detection efficiency. The wiring used to connect an individual detector to the pre-amplifier leads to parasitic capacitance, inductance and resistance and therefore increases the over all noise to signal ratio.